Scintillation camera systems for use in positron imaging have been employed for purposes of medical diagnosis to only a limited extent in the field of nuclear medicine. An early description of a positron imaging device employing scintillation crystals was disclosed in a report by Hal O. Anger printed for the U.S. Atomic Energy Commission entitled "Scintillation and Positron Cameras" (UCRL-8640, Aug. 12, 1959). The underlying concept of a scintillation camera is set forth in U.S. Pat. No. 3,011,057. The positron imaging scintillation camera system described in the aforesaid report is a particular application of a scintillation camera system in which alterations and modifications to a basic scintillation camera system are employed to enable detection of positron events. Investigation was undertaken into possible techniques and instrumentation which might be advantageously employed in scintillation camera systems for positron imaging. The results of some of these investigations are reported in: Anger, H. O.: Radioisotope Cameras: Instrumentation in Nuclear Medicine, Vol. 1, Hine, G. J. editor, New York, Academic Press, 1967; Brownell, G. L., Burham, C. A.: Recent Developments in Positron Scintigraphy: Instrumentation in Nuclear Medicine, Vol. 2, Hine, G. J. Sorenson, J. A., editors, New York, Academic Press, 1973; and Kenny, P. J.: Spatial Resolution and Count Rate Capacity of a Positron Camera: Some Experimental and Theoretical Considerations: International Journal of Applied Radiation and Isotopes, Vol. 22, Permagon Press, pp. 21-28, 1971. The results of these investigations have produced the conclusions that the useful count rate of a positron camera is only a small fraction of the actual count rate. This conclusion has been based largely on the proposition that only the photopeak produced by the two 511 KeV gamma rays emitted after the decay of a positron may be used to register a radioactive distribution of positron events within a subject of interest. In a conventional positron imaging system, two scintillation detectors are positioned on opposite sides of a subject of interest. Typically, the subject under investigation is an organ of interest of a living human being lying in a prone position. One of the scintillation detectors is positioned above the subject while the other is positioned beneath. Each of the scintillation detectors is comprised of a disc of sodium iodide, typically one-half an inch in thickness, viewed by an array of photomultiplier tubes which detect scintillations occurring in the crystal and generate electrical pulses in response thereto. The electrical pulses from both detectors are used to calculate the location of the activity for a particular plane through the subject of interest. In contradistinction to low energy gamma rays, where approximately 90% of all gamma rays contribute a count to the photopeak, only about 17 % of the positron gamma rays are in the 511 KeV photopeak. About 18% of the positron gamma rays undergo Compton scatter in the crystal followed by escape of the secondary photon, and the rest penetrate a one-half inch thick sodium iodide scintillation crystal without any kind of interaction. The degree to which positron annihilations yield useful information is further reduced by the requirement in a positron imaging system for coincident detection of the two gamma rays emitted by the positron. When a positron is annihilated in the subject, two gamma rays of 511 KeV energy are produced which travel in approximately opposite directions. Since only a relatively small fraction of the positron annihilations are detected, it is particularly important to distinguish the detected positron events from background radiation. This is done through the requirement for coincidence. In addition, the requirement for coincidence is also necessary when it is desired to locate and distinguish between planes at which positron events occur. A system for accomplishing this latter object is disclosed in U.S. Pat. No. 3,573,458.